X-ray systems are generally used to irradiate a body with X-ray radiation by using an X-ray emitter and to display the penetration through the body by using, for example, a fluorescent screen or an image amplifier. The images are either visualized on suitable film material or phosphorous plates; or by using electronic sensors such as CCDs.
X-ray emitters generally contain an X-ray tube. In its simplest form, the latter includes a cathode and an anode which sit in a vacuum within a sealed glass- or metal body. The cathode emits electrons which are accelerated by high voltage toward the anode and penetrate the anode material. In the process, they are decelerated and mainly generate characteristic X-ray radiation and X-ray bremsstrahlung.
Characteristic X-ray radiation is generated by the energetic electrons accelerated in the X-ray tube knocking out electrons from the innermost shells of the atoms of the anode material in the anode. Electrons from higher energy levels or free electrons “jump” into these gaps which are created. The energy liberated in this process is emitted in the form of discrete energy quanta which are typical for the material.
By contrast, X-ray bremsstrahlung is generated by the deceleration of the electrons when passing through the metal of the anode since every accelerated electric charge generates electromagnetic radiation. The wavelength of the radiation depends on the value of the acceleration such that a higher acceleration voltage or anode voltage generates more X-ray radiation with more energetic quanta. Accordingly, the maximum energy of the bremsstrahlung spectrum is the energy at which the entire kinetic energy of the electron is transferred to a single quantum. This limiting energy only depends on the anode voltage and is independent of the anode material.
In the case of medical X-ray examinations, different radiation energies are required for different regions of the body in order to penetrate through tissues of different densities, such as fatty tissue or bone. As explained above, the voltage supplied to the X-ray tube is decisive in this case. Depending on the desired information to be supplied by the image, the tube voltage is selected between, for example, 38 kV and 120 kV. In the case of low energies, a lot of radiation is absorbed by the tissue; this makes very fine differences in the tissue visible on the X-ray film. By contrast, high-energy radiation penetrates through tissue and materials much more easily and contrast differences are strongly watered down.
In the case of a conventional X-ray method, the object to be imaged is irradiated by an X-ray source and imaged on X-ray film. A projection of the volume on a plane is generated. In this projection, information relating to the third dimension of the irradiated body is lost to a great extent. The reason for this is that retrospectively it is no longer possible to distinguish between attenuation, visible in the X-ray image, being caused by a more dense material or by a larger slice thickness.
This problem is solved by computed tomography (CT), in which many X-ray images of the object are generated from the most diverse directions, and volume information is subsequently reconstructed from these many images. In general, these 3D reconstructions are composed of individual slices which run across the object. This affords the possibility of determining a density for each volume element of the object.
Since computed tomography requires a multiplicity of X-ray images to be recorded within a very short timeframe, CT detectors are required which can directly supply the X-ray image to a data processing unit in digital form. For this purpose, electronic detectors are usually used, e.g. solid-state detectors. In order to spatially resolve the X-ray image, these detectors usually comprise individual X-ray sensors which are arranged like pixels.
The individual sensors of the detector can have very different designs but generally operate on the basis of the inner photoelectric effect. Here, an electric pulse is generated when an X-ray quanta impinges on the sensor, with the pulse amplitude being characteristic of the energy of the quantum. The pixel-like arrangement of the sensors thus in principle affords the possibility of resolving, spatially and energetically, every single incident X-ray quantum.
In order to count the incident X-ray quanta in the connected data processing device, the pulses of the sensors can be counted by a trigger circuit. Here, an energy threshold is prescribed and a count signal is emitted when the pulse amplitude of the sensor corresponding to this energy is exceeded (single pulse triggering). For conventional imaging, this threshold is selected such that it for example is in the range from 15 keV to 35 keV. In the case of dual-energy imaging, provision is made for a further threshold, e.g. in the range from 50 keV to 80 keV.
Due to the high rates of X-ray quanta impinging on the detector occurring in computed tomography, the in principle appealing concept of a detector counting individual quanta constitutes a problem which is difficult to solve. One of the critical points is the finite pulse duration which in typical sensors (e.g. CdTe or CdZnTe) is e.g. approximately 10 ns (full width at half maximum). In conjunction with a required electronic pulse formation this results in effective pulses of the order of e.g. approximately 30 ns which should be registered individually.
In the case of temporally equidistant incident pulses, this alone would limit the maximum measurable rate to approximately 33 MHz per pixel, which in the case of realistic pixel sizes with an edge length of e.g. approximately 200 μm corresponds to a maximum X-ray quanta flux of approximately 825 MHz/mm2. However, since the temporal arrival of the pulses is in fact subject to Poisson statistics, the probability of pulses at least partly overlapping with one or more further pulses is already 26% in the case of (on average) 33 million quanta per pixel and second. In the case of maximal fluxes occurring in computed tomography these days, which are approximately 2 GHz/mm2, this probability actually increases to over 60%. This implies that, despite using e.g. shape filters in the beam path, detector channels which have no or little attenuation, e.g. at the edge of the objects to be examined, are no longer able to resolve individual pulses since the pulses of the incident quanta superpose, possibly even superpose a number of times.
In the case of such a multiple superposition, the pulse amplitude no longer sinks below the level of the predetermined threshold energies of the triggers after every pulse. These then trigger fewer and fewer signals and therefore the signal rate can no longer be unambiguously assigned to the actual quanta flux; this is referred to as the onset of paralysis. In the case of even higher quanta fluxes, it is more and more seldom for the pulse amplitude to drop below the level of the predetermined threshold energies of the triggers since the pulse amplitude is continuously increased by superposing pulses. The trigger then no longer triggers a signal; the signal rate tends to zero. In regions of relatively strong superposition, i.e. at the edge of the observed objects or in the air, for example, this can result in data having to be discarded.